Image-forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.
According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse) while the magnetic field extends perpendicular to the z-axis, so that the magnetization performs a precessional motion about the z-axis. The precessional motion describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°).
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of receiving RF coils which are arranged and oriented within an examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneities) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing). The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.
In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The signal data obtained via the receiving coils corresponds to the spatial frequency domain and is called k-space data. The k-space data usually includes multiple lines acquired with different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to an MR image, e.g., by means of Fourier transformation.
Cardiac interventional MR imaging is a promising tool in which accurate localization of an interventional instrument with excellent soft tissue contrast can be combined. Moreover, functional information from the heart can be obtained by means of appropriate MR imaging techniques. The combination of MR imaging with tracking of interventional instruments is especially advantageous for therapeutic applications that require therapy monitoring, like, e.g., MR electrophysiology interventions. However, cardiac MR imaging is associated with a compromise between spatial resolution, scan time and signal-to-noise ratio (SNR). Therefore effective motion compensation is of utmost importance. Acquisition of sufficient MR data for reconstruction of an image takes a finite period of time. Motion of the object to be imaged, like the beating motion of the heart in combination with the respiratory motion of the patient, during that finite acquisition time typically results in motion artifacts in the respective reconstructed MR image. The acquisition time can be reduced to a very small extend only, when a given resolution of the MR image is specified. In dynamic MR imaging scans, as required for therapy monitoring, the motion of the examined object during data acquisition leads to different kinds of blurring, mispositioning and deformation artifacts. Prospective motion correction techniques, such as the so called navigator technique or PACE, have been developed to overcome problems with respect to motion by prospectively adjusting the imaging parameters, i.e. the parameters of the imaging sequence used for MR signal acquisition, which define the location and orientation of the field of view (FOV) within the imaging volume. In the navigator technique, a MR data set is acquired from a pencil-shaped volume (navigator beam) that crosses the diaphragm of the examined patient. The volume is interactively placed in such a way that the position of the diaphragm can be reconstructed from the acquired MR data set and used for motion correction of the FOV in real time. The navigator technique is primarily used for minimizing the effects of breathing motion in cardiac examinations. Opposed to the navigator technique, which requires a navigator beam to detect motion differences, the above-mentioned PACE technique uses previously acquired dynamic images to prospectively adjust the imaging parameters on the time scale of successive dynamic scans. Moreover, it is known to apply ECG-based gating for the purpose of synchronization of the image acquisition with the beating motion of the heart, thereby reducing motion artifacts due to cardiac cycling.
The known approaches of motion compensation disadvantageously require an increased scan time due to the decreased scan duty cycle. Moreover, the above-mentioned navigator technique requires complex scan planning.
On the other hand, it has recently been shown that MR imaging is capable of visualizing the effect of a cardiac electrophysiology ablation shortly after the ablation, wherein it was demonstrated that ablation-related physiologic changes can be identified by means MR imaging in-situ. However, presently limitations exist with respect to image quality due to limited SNR and motion artifacts.